Classifications

• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/72—Signal processing specially adapted for physiological signals or for diagnostic purposes
  • A61B5/7271—Specific aspects of physiological measurement analysis
  • A61B5/7278—Artificial waveform generation or derivation, e.g. synthesising signals from measured signals
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
  • G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
  • G01N21/17—Systems in which incident light is modified in accordance with the properties of the material investigated
  • G01N21/21—Polarisation-affecting properties
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/0033—Features or image-related aspects of imaging apparatus, e.g. for MRI, optical tomography or impedance tomography apparatus; Arrangements of imaging apparatus in a room
  • A61B5/0035—Features or image-related aspects of imaging apparatus, e.g. for MRI, optical tomography or impedance tomography apparatus; Arrangements of imaging apparatus in a room adapted for acquisition of images from more than one imaging mode, e.g. combining MRI and optical tomography
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/0059—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/0059—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence
  • A61B5/0071—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence by measuring fluorescence emission
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/0059—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence
  • A61B5/0075—Measuring for diagnostic purposes; Identification of persons using light, e.g. diagnosis by transillumination, diascopy, fluorescence by spectroscopy, i.e. measuring spectra, e.g. Raman spectroscopy, infrared absorption spectroscopy
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/145—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue
  • A61B5/14546—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue for measuring analytes not otherwise provided for, e.g. ions, cytochromes
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/145—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue
  • A61B5/1455—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using optical sensors, e.g. spectral photometrical oximeters
  • A61B5/14558—Measuring characteristics of blood in vivo, e.g. gas concentration or pH-value ; Measuring characteristics of body fluids or tissues, e.g. interstitial fluid or cerebral tissue using optical sensors, e.g. spectral photometrical oximeters by polarisation
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/44—Detecting, measuring or recording for evaluating the integumentary system, e.g. skin, hair or nails
  • A61B5/441—Skin evaluation, e.g. for skin disorder diagnosis
  • A61B5/443—Evaluating skin constituents, e.g. elastin, melanin, water
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B5/00—Measuring for diagnostic purposes; Identification of persons
  • A61B5/44—Detecting, measuring or recording for evaluating the integumentary system, e.g. skin, hair or nails
  • A61B5/441—Skin evaluation, e.g. for skin disorder diagnosis
  • A61B5/444—Evaluating skin marks, e.g. mole, nevi, tumour, scar
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01N—INVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
  • G01N21/00—Investigating or analysing materials by the use of optical means, i.e. using sub-millimetre waves, infrared, visible or ultraviolet light
  • G01N21/17—Systems in which incident light is modified in accordance with the properties of the material investigated
  • G01N21/25—Colour; Spectral properties, i.e. comparison of effect of material on the light at two or more different wavelengths or wavelength bands
  • G01N21/31—Investigating relative effect of material at wavelengths characteristic of specific elements or molecules, e.g. atomic absorption spectrometry

Definitions

• Melanoma is a serious and challenging disease. It is an increasingly lethal form of skin cancer, especially when detected in later stages. Melanoma risk during a lifetime increased from 1 :1500 in 1935 to 1 :58 in 2009, and is still the fastest growing cancer both in the U.S. and worldwide.
• the National Cancer Center has estimated that 76,250 patients will be diagnosed with melanoma of the skin in 2012 and that 9,180, or more than one patient per hour, will die.
• the present common standard in melanoma patient care is a dermatologists' visual examination, such as the ABCDE procedure or revised 7-point checklist in which the practitioner looks for abnormalities in shape, size and color.
• MoleMateTM is a 4-color, light emitting diode (“LED”) based non-invasive melanoma screening device that employs
• SIA scans are used to gather information about a patient's suspicious moles and lesions by imaging pigment, collagen, and blood directly under the mole or lesion.
• blackbox methods based on statistical classifiers.
• Mela Sciences Inc. is a registered trademark of Mela Sciences Inc.
• Verisante Aura is claimed as a trademark by The BC Cancer Agency and the University of British Columbia
• blackbox approach assumes there is an optical signature difference between normal and cancerous tissues and addresses differentiation between these tissue states by using statistical classifiers and training-based discrimination functions.
• Unfortunately many systems employing these methods have shown reductions in performance as the studies move from smaller to larger populations.
• a telling example is the specificity reduction in the MelaFind® device from 84% reported in 2001 compared to results in the 9.5% to 11% range in 2011.
• the MelaFind® device data shows unavoidable rates of false-positives and false-negatives.
• One method that does provide depth information is high resolution confocal microscopy such as that performed by the VivaScope confocal microscope (a registered trademark of Caliber Imaging & Diagnostics, Rochester, NY). It takes a microscopic image of a shallow depth of skin lesion ( ⁇ 700 ⁇ ) and small field of view (FOV ⁇ 1 mm x lmm), which is then analyzed by a dermal pathologist to detect melanoma or other skin cancers. These devices are very expensive, and the interpretation of the information requires the skills of a pathologist. There is still an unmet a need for a device that provides simple-to-interpret depth information about a melanoma lesion at a reasonable cost.
• Real time computation (30 ms to 1000 ms) is ideal for extracting high resolution skin chromophore two-dimensional maps from three- dimensional spectral image stacks with millions of voxels.
• These rapid quantification algorithms range from ratiometric calculations of skin reflectance maps at various wavelengths to Beer-Lambert, Attas, M., Hewo, M., Payette, J., Posthumus, T., Sowa, M., et al. Visualization of cutaneous hemoglobin oxygeneation and skin hydration using near-infrared spectroscopic imaging., Skin Res Techno!, 7, 238- 245 (2001) or two-flux ubelka-Munk models (up to few minutes) for homogenous turbid media, Anderson, R.
• Levenberg-Marquardt [Zonios, G., Bykowski, J., & Kollias, N. Skin melanin, hemoglobin, and light scattering properties can be quantitatively assessed in vivo using diffuse reflectance spectroscopy., J Invest Dermatol, 117(6), 1452-1457 (2001), to fit the measured spectral signature with the estimated spectral signature derived from the related forward model.
• Non-invasive, dynamic fluorescence imaging of human melanomas reveals that targeted inhibition of bFGF and FGFR-1 blocks tumor growth by inducing melanoma cell apoptosis.
• Mol Med, 8, 103-112 (2002), Pfaff-Smith, A., Kirkwood, J.M., Edington, H.D., Jukic, D.M., Farkas, D.L. et al. Fluorescence imaging analysis of upstream regulators and downstream targets of STAT3 in melanoma precursor lesions obtained from patients before and after systemic low-dose interferon-a treatment., Mol Imaging, 2, 65-73 (2003).
• a method and system are provided for characterizing a portion of biological tissue.
• a disclosed system comprises a source of multiple wavelength light is configured to illuminate said tissue with a temporal sequence of different wavelengths to produce corresponding images of said hyperspectral image sets.
• the model of tissue may comprise a theoretically generated model or an empirically generated model.
• the empirically generated model is based on measurements of illuminated normal tissue or measurements of an illuminated tissue phantom.
• the characteristics of the abnormal portion of the tissue may be produced by solving an inverse problem based on the model, starting with the measurements of intensity at a plurality of wavelengths and a plurality of polarizations and modifying estimation parameters of the model to produce a solution to the problem that substantially matches the characteristics of the tissue.
• the concentration of superficial melanin is separated in order to quantify the deep melanin relative
• Fig. 1 is a schematic representation of a cross sectional view of biological tissue with two different anomalies at the same depth and graphs illustrating the corresponding degree of linear polarization spectra variation for the respective anomalies.
• Fig. 2 is a schematic representation of a cross sectional view of biological tissue with two identical anomalies at different depths, and graphs illustrating the corresponding degree of linear polarization spectra variation for the respective anomalies.
• Fig. 3 is a schematic representation of a cross sectional vie of a tissue sample showing different depths of illuminating light penetration at different illumination wavelengths.
• Fig. 4 is a diagramatic representation of a method for extracting depth
• Fig. 5 is a cross sectional image of skin showing melanin in superficial and deep layers of the skin.
• Fig. 6(a) is an illustration of a uniform melanin distribution in skin.
• Fig. 6(b) is an illustration of a non-uniform melanin distribution in skin.
• Fig. 7 shows the top view of an example reflectance image of skin exhibiting a feature characteristic of vitiligo.
• Fig. 8 illustrates the measurements of degree of linear polarization spectra for an example of skin exhibiting a feature characteristic of vitiligo.
• Fig. 9 shows the top view of an example reflectance image of skin exhibiting a feature characteristic of mole condition.
• Fig. 10 illustrates the measurements of degree of linear polarization spectra for an example of skin exhibiting a feature characteristic of mole.
• Fig.11 is a diagramatic representation of a method for creating a three
• Fig. 12 is a block diagram representation of a system for capturing tissue data using multiple optical modes.
• Fig. 13 is a diagramatic representation of a method for creating a three
• Fig. 14 is a diagramatic representation of a method for creating a three
• Fig. 15 is a block diagram of a system for implementing hyperspectral, polarization distinguishing optical measurements as described herein.
• Fig. 16 is an illustration of hyperspectral datacubes produced by the system of Fig. 15.
• Fig. 17(a) is a colored image of a patch of a subject's skin having melanocytic nevus, illuminated with linearly polarized light having a uniform intensity spectrum from 475nm to 825nm and acquired through a parallel linear polarizer.
• Fig. 17(b) is a colored image of the skin patch of Fig. 17(a), illuminated as in Fig. 17(a), but acquired through a crossed linear polarizer.
• Fig. 17(c) is a graph of crossed-polarization optical density as a function of wavelength for a central (melanocytic nevus core) region, a boundary (halo) region, and a surrounding (normal skin) region of the skin patch.
• Fig. 17(d) is a graph of polarized attenuation as a function of wavelength a central (nevus core) region, a boundary (halo) region, and a surrounding (normal skin) region of the skin patch.
• Fig. 17(e) is a colored image of a patch of a subject's skin exhibiting vitiligo, illuminated with linearly polarized light having a uniform intensity spectrum from 475nm to 825nm and acquired through a parallel linear polarizer.
• Fig. 17(f) is a colored image of the skin patch of Fig. 14(a), illuminated as in Fig. 14(a), but acquired through a crossed linear polarizer.
• Fig. 17(g) is a graph of crossed-polarization optical density as a function of wavelength for the central region (little or no melanin), the boundary region (some melanin), and the surrounding region (high concentration of melanin) of the skin patch.
• Fig. 17(h) is a graph of polarized attenuation as a function of wavelength a central (vitiligo) region, a boundary (halo) region, and a surrounding (normal skin) region of the skin specimen.
• Fig. 18(a) is a concatenated sequence of thirteen colored images of a patch of skin on the dorsal side of a finger of a subject experiencing normal blood flow, wherein the images have been taken in thirty-second intervals.
• Fig. 18(b) is a concatenated sequence of thirteen colored images of the patch of skin of Fig. 18(a), wherein the images have been taken in thirty-second intervals and the subject's blood flow has been occluded by a cuff around the finger for a one hundred fifty second interval after the beginning and before the end of that sequence.
• Fig. 18(c) is a graph of crossed-polarization optical density as a function of wavelength for region I of Fig. 18(b) (before occlusion), region II of Fig. 18(b) (during occlusion), and region III of Fig. 18(b) (after occlusion).
• Fig. 18(d) is a graph of the polarization attenuation as a function of wavelength for region I of Fig. 18(b) (before occlusion), region II of Fig. 18(b) (during occlusion), and region III of Fig. 18(b) (after occlusion).
• Fig. 18(e) is a concatenated sequence of thirteen colored images of a patch of skin on the volar side of a finger of a subject experiencing normal blood flow, wherein the images have been taken in thirty-second intervals.
• Fig. 18(f) is a concatenated sequence of thirteen colored images of the patch of skin of Fig. 18(e), wherein the images have been taken in thirty-second intervals and the subject's blood flow has been occluded by a cuff around the finger for a one hundred fifty second interval after the beginning and before the end of that sequence.
• Fig. 18(g) is a graph of crossed-polarization optical density as a function of wavelength for region I of Fig. 16(f) (before occlusion), region II of Fig. 16(f) (during occlusion), and region III of Fig. 16(f) (after occlusion).
• Fig. 19 is a flow chart of a process for estimating the quantity of hemoglobin to characterize a tissue anomaly.
• Fig. 20(a) is a composite (red-green-blue) image of the skin patch of Figs. 17(a) and 17(b) having a melanocytic nevus.
• Fig. 20(b) is a chromophore map of the melanin in the skin patch of Fig. 20(a) derived from the optical density spectra in cross-polarization mode.
• Fig. 20(c) is a sequence of chromophore maps of oxy-hemoglobin (oHb), deoxy- hemoglobin (Hb), total hemoglobin (tHb) and an oxygen saturation parameter (OSP) in the skin patch of Fig. 20(a) illustrating how a high melanin concentration is conducive to mis-estimation of hemoglobin concentrations.
• oHb oxy-hemoglobin
• Hb deoxy- hemoglobin
• tHb total hemoglobin
• OSP oxygen saturation parameter
• Fig. 20(d) is a sequence of chromophore maps of oxy-hemoglobin (oHb), deoxy- hemoglobin (Hb), total hemoglobin (tHb) and an oxygen saturation parameter (OSP) in the skin patch of Fig. 20(a) derived from a two-chromophore model to correct for the presence of hemoglobin.
• oHb oxy-hemoglobin
• Hb deoxy- hemoglobin
• tHb total hemoglobin
• OSP oxygen saturation parameter
• Fig. 20(e) is a composite (red-green-blue) image of the skin patch of Figs. 17(e) and 17(f) exhibiting vitiligo.
• Fig. 20(f) is a chromophore map of the melanin in the skin patch of Fig. 20(e) derived from the optical density spectra in cross-polarization mode.
• Fig. 20(g) is a sequence of chromophore maps of oxy-hemoglobin (oHb), deoxy- hemoglobin (Hb), total hemoglobin (tHb) and an oxygen saturation parameter (OSP) in the skin patch of Fig. 20(e) illustrating how a high melanin concentration is conducive to mis-estimation of hemoglobin concentrations.
• oHb oxy-hemoglobin
• Hb deoxy- hemoglobin
• tHb total hemoglobin
• OSP oxygen saturation parameter
• Fig. 20(h) is a sequence of chromophore maps of oxy-hemoglobin (oHb), deoxy- hemoglobin (Hb), total hemoglobin (tHb) and an oxygen saturation parameter (OSP) in the skin patch of Fig. 20(e) derived from a two-chromophore model [and hyperspectral polarized images] to correct for the presence of hemoglobin.
• oHb oxy-hemoglobin
• Hb deoxy- hemoglobin
• tHb total hemoglobin
• OSP oxygen saturation parameter
• 21(a) is a Relative molar absorptivity of Oxy-hemoglobin (oHb), deoxy-hemoglobin (Hb), total hemoglobin (tHb), melanin and oxygen saturation (OSP) maps with corresponding color cross-polarized image of dorsal finger during finger cuff occlusion
• oHb Oxy-hemoglobin
• Hb deoxy-hemoglobin
• tHb total hemoglobin
• OSP melanin and oxygen saturation
• Fig. 21(b) is a Relative molar absorptivity of Oxy-hemoglobin (oHb), deoxy-hemoglobin (Hb), total hemoglobin (tHb), melanin and oxygen saturation (OSP) maps with corresponding color cross-polarized image of volar finger during finger cuff occlusion
• oHb Oxy-hemoglobin
• Hb deoxy-hemoglobin
• tHb total hemoglobin
• OSP melanin and oxygen saturation
• This disclosure presents preferred embodiments of a system and method that enable the production of a three dimensional map that provides volumetric information about tissue biology from analysis of multimode hyperspectral data cubes.
• Data cubes are sets of images taken under multiple modalities which can be analyzed.
• the images for the datasets are captured by a multimode imaging system such as the SkinSpect multimode imaging system developed by Spectral Molecular Imaging, Beverly Hills, CA, which combines hyperspectral, polarization, reflection, scattering, fluorescence and bio- fluorescence imaging modalities.
• the three-dimensional optical map created from this data provides information to the physicians that helps to diagnose tissue abnormalities with higher precision than with imaging data sets having fewer modalities and
• Hyperspectral imaging is the capture of a sequence of images of a
• target such as tissue at multiple wavelengths of light that include wavelengths outside the visible spectrum, where each image contains data indicative of theproperties of remitted light in a specific narrow wavelength band.
• "Remitted” light includes reflected and scattered light, and fluorescent, luminescent and bio-luminsecent light produced in response to illumination light.
• the narrow wavelength band can be created by controlling the light illuminating the target, or it can be created by controlling the light emanating from the target.
• the sequence of images captured for a hyperspectral image provides reflectance, scattering or other emission data at multiple wavelengths sufficient to reconstruct with reasonable accuracy the reflectance, scattering or other emission spectrum of the target at each image pixel of the target.
• these spectra will have wavelength data intervals ranging between 1 nm and 50 nm but these intervals may be smaller or larger depending on the nature of the spectrum and the needs of the analysis.
• Polarization imaging is the capture of a sequence of images of a target such as tissue, where each image contains data indicative of the polarization properties of the target.
• Light reflected scattered or otherwise remitted from a tissue can have its polarization properties modified by its passage into or out of a tissue.
• Polarization images can be created by filtering or otherwise controlling polarization of the light illuminating or remitted from a tissue, or both, and capturing images of light with particular polarization properties.
• a computer is used in connection with the acquisition and processing of acquired data to generate an enhanced map, or multi-dimensional data base, of the structural characteristics of the tissue being measured.
• the following reports provided to a clinician automatically and quantitatively: (1) both ABCDE and modified 7-point checklists, (2) three dimensional maps of tissue composition allowing both area and cross-sectional views that can selectively show melanocyte progression, hemoglobin distribution, collagen and elastin abnormalities, and angiogenesis, and (3) surface topology of skin lesions including spatial analysis reports.
• Obtaining the depth of a melanoma lesion is of cardinal importance in successful early diagnosis.
• the invention requires much more comprehensive measurement data than is used in other methods and apparatuses.
• three dimensional distributions of biological features in tissue are obtainable.
• Melanoma surface lesions are very easy to remove. Lesions with a depth greater than one millimeter quickly become lethal.
• the embodiments disclosed herein use more comprehensive measurement data, and then apply this data to well- developed and understood physiological tissue models to provide quantitative measures of the spatial distribution of biological features in tissue.
• the method and system disclosed herein employ data from a combination of optical techniques, including diffuse reflectance spectroscopy, the polarization of light remitted from the tissue, Mie-scattering analysis and tissue fluorescence, luminescence or bio-luminescence in an imaging mode to produce maps of the distribution of tissue features from the surface to depths of up to 2 mm.
• Polarization filtered fluorescence imaging data is used to determine fluorescence anisotropy analytically to quantify tissue features such as collagen and elastin distribution.
• Diffuse reflectance hyperspectral imaging is used to quantify hemoglobin, melanin, water and fat distribution, as well as scattering properties of tissue, which can provide information about growth characteristics and cell proliferation.
• the multimodal nature of the imaging data allows extraction of information to apply to inverse models of tissue optical properties. This method can detect, correct and compensate for data analysis uncertainties that straight spectral imaging or multi-wavelength imaging cannot.
• Model-based feature extraction from image data eliminates much of the measurement variability that can plague statistical methods, especially when correlated against associated features or features from neighboring voxels in the image data sets.
• the direct linkage to underlying tissue characteristics provides the biological plausibility that many clinicians require before adopting a technology. This biological plausibility also makes the method and system more easily testable, using tissue phantoms and appropriate standards to verify accuracy of quantification and ongoing system
• a new method and an apparatus are disclosed that use two depth-sensitive techniques: polarization and hyperspectral imaging, to accurately determine the spatial distribution of melanin and hemoglobin oxygenation in a skin lesion.
• the method and apparatus accurately separate the contribution of superficial melanin in order to quantify U 2014/014330
• the deep melanin is primarily found in melanosomes migrating to the skin superficial layer as a part of normal epidermal replacement. Lin, J. Y., & Fisher, D. E. Melanocyte biology and skin pigmentation Nature, 445, 843-850 (2007).
• the deep melanin is primarily associated with the melanocytes found on the basal layer that separates the epidermis and dermis layers.
• a linearly polarized, multi-wavelength light source is used to illuminate the skin while both parallel and perpendicular polarization images of the remitted light are recorded simultaneously by two cameras.
• This effect is illustrated herein using skin with a melanocytic nevus (high melanin) and skin with vitiligo (low melanin) as well as skin under the influence of venous occlusion (changing hemoglobin) to demonstrate the effectiveness of this method for accurately distinguishing and quantifying hemoglobin and melanin distributions.
• the method and system for in vivo tissue characterization disclosed herein employ illumination of tissue with hyperspectral, polarized light and spatial
• Remitted light is intended to refer to light that is specularly reflected, diffusely reflected or back scattered, or light remitted as fluorescence, luminescence or bio-luminescence, or combinations of the foregoing. Spatial measurements of intensity as a function of wavelength and relative polarization of remitted light have been found to enable
• living tissue is illuminated with a spectrum light preferably in the visible and near-infrared spectrum, typically having material wavelengths ranging from about 400 nanometers to about 1000 nanometers, though the ends of the spectrum used ordinarily are neither sharp nor critically important.
• the wavelength spectrum over this range may be continuous or discontinuous, even discrete, depending on the particular need.
• the illumination light is polarized in a known way. Ordinarily, linear polarization would be used, but other polarizations such as circular or elliptical might be used without departing from the principles of the invention.
• the intensity of light remitted from the skin in response to the illumination is measured in a hyperspectral measurement space which is ordinarily two-dimensional. However, it is to be understood that one-dimensional or three-dimensional
• the polarization of remitted light is indicative of the physiologic character of the tissue remitting the light.
• the polarization of the remitted light relative to the polarization of the illumination light is indicative of the tissue character.
• This is expressed as the degree of polarization, in particular, the degree of known input polarization in the remitted light.
• the input polarization would be linear, and the degree of linear polarization of the remitted light would be measured.
• the degree of linear polarization of remitted light DLP(X) at a given point in measurement space may be expressed as a function of wavelength as:
• ⁇ is the wavelength of light
• ⁇ ⁇ ( ⁇ ) is the intensity of linearly polarized remitted light parallel to the input polarization at wavelength ⁇ ;
• ⁇ ⁇ ( ⁇ ) is the intensity of linearly polarized remitted light perpendicular to the input polarization at wavelength ⁇ . More generically it is to be recognized that circularly or elliptically-polarized light might also be used and the degree of polarization would compare the orthogonally polarized light in the remitted light with the remitted light having the input polarization.
• Tissue fluorescence, luminescence and bio-luminescence remitted in response to input light may also characterize anomalies in the tissue.
• a comparison of the intensity of remitted fluorescence, luminescence or bio-luminescence light with the illumination can be used for this purpose.
• a three-dimensional model of the tissue comprising an array of individual tissue-characteristic three-dimensional voxels may be produced.
• the embodiments disclosed herein can be basically understood by considering the task of locating, characterizing and distinguishing two different anomalies in tissue, as explained hereafter.
• FIG. 1 A schematic representation of a cross section of tissue 1 with two types of anomalies is shown in Fig. 1.
• One anomaly 2 has lower optical attenuation and the other anomaly 3 has higher optical attenuation, both relative to the surrounding normal tissue 1.
• Both anomalies are illuminated when the tissue is illuminated by polarized light 4.
• the light remitted from tissue 1 is filtered to selectively pass different polarizations of light.
• the filtered polarized light can be detected by a photo-detector such as a photodiode, charge coupled device, or similar light measurement device to obtain a 2D or 3D image data set.
• the photo-detector may comprise a single point measurement system or may comprise an array of detectors such as an image capture device.
• a single point measurement system may also optically or mechanically scan to capture an array of measurements.
• the remitted light intensity is measured by the detector and the intensity information is stored for at least two polarization states.
• the different linear polarizations of the measurement data are captured using cross 5 and parallel 6 of the polarizing filter in the detection path with respect to the orientation of the polarizing filter in the illumination path. These two polarization images are captured using in at least two different wavelengths of illumination light.
• multiple wavelengths of illumination are used, for example greater than 30 wavebands.
• the data captured can form three dimensional data cubes for both parallel and cross polarization states.
• the data is preferably captured is in the form of an image and the three dimensional data is a hyperspectral image cube containing x and y spatial coordinates as well as the intensity of remitted light from tissue at each wavelength of illumination and for each polarization state.
• the evaluation of the optical attenuation in the anomalies can be performed by analysis methods incorporating the diffuse reflectance wavelength dependence of the degree of linear polarization parameter
• Ip and ⁇ are reflectance intensity at parallel and cross (orthogonal) linear polarization modes, and ⁇ is the wavelength of illumination.
• the anomaly with higher optical attenuation 3 than the surrounding normal tissue experiences less cross polarized signal 5, which leads to the higher degree of polarization value 10 in graph 120.
• the reflectance intensity from the anomaly with lower optical attenuation 2 provides higher cross polarization detected light intensity 5; therefore, its translation to degree of polarization spectra 110 shows more attenuation 9 compared to surrounding normal tissue DLP spectra 8.
• Fig. 2 there is shown a schematic representation of a cross section of tissue 1 with two identical anomalies that produce higher optical attenuation relative to the surrounding normal tissue.
• One the anomalies 21 is deeper beneath the tissue surface.
• the second anomaly 3 is close to the tissue surface.
• both anomalies are illuminated by polarized light 4.
• the light remitted from tissue 1 can be filtered to selectively pass different polarizations of light.
• the filtered polarized light can be detected by a photo detector such as a photodiode, charge coupled device, or similar light measurement device.
• the photo-detector may comprise a single point measurement system or may comprise an array of detectors such as an image capture device.
• a single point measurement system may also optically or mechanically scan to capture an array of measurements.
• the remitted light intensity is measured by the detector
• the intensity information is stored for at least two polarization states.
• the different linear polarizations of the measurement data are captured using parallel 6 and cross 5 orientations of the polarizing filter in the detection path with respect to the orientation of the polarizing filter in the illumination path. These two polarization images are captured in at least two different wavelengths of illumination light.
• the data captured can form three dimensional data cubes for both parallel and cross polarization states.
• the data is captured preferably in the form of an image and the three dimensional data is a hyperspectral image cube containing x and y spatial coordinates as well as the intensity of remitted light from tissue at each wavelength of illumination and for each polarization state.
• the evaluation of the optical attenuation in the anomalies can be performed by analysis methods incorporating the diffuse reflectance wavelength dependence of degree of linear polarization parameter
• DLP( ) ( ⁇ ⁇ ( ⁇ ) - ⁇ ( ⁇ ))/( ⁇ ⁇ ( ⁇ ) + ⁇ ⁇ ( ⁇ ))
• I P and I x are reflectance intensity at parallel and perpendicular (cross) linear polarization modes and ⁇ is the wavelength of illumination.
• the anomaly at greater depth 210 experiences lower cross polarized signal 5 in longer wavelength ranges which leads to the higher degree of polarization value 22 in graph 220.
• the DLP spectral signature in 1 10, 120, 210, and 220 illustrates the power of DLP spectral signature to differentiate the effect of signal changes due to the optical attenuation of an anomaly and the effect of signal changes due to the depth of an anomaly.
• FIG. 4 the schematic representation of a cross section of a tissue sample 310 illustrate that DLP 320 values calculated from measurements at different wavelengths are representative of tissue characteristics at different depths 330 within the tissue. The longer the wavelengths of illumination, the deeper in the tissue the information comes from. This DLP information can be correlated to the depth within the tissue to create three dimensional maps of tissue optical properties.
• Method 400 comprises first at 410 receiving measurements of intensity at each spatial coordinate from hyperspectral data- cubes of tissue sample measurements in both linear polarization modes (i.e. parallel polarized light intensity ⁇ ⁇ ( ⁇ ) and cross polarized light intensity I x ( ⁇ )). Then, at 420, based on those measurements the method features further comprises calculating the degree of linear polarization DLP ( ⁇ ).
• the resulting depth resolved optical properties of tissue 530 can be used to identify the tissue composition (such as melanin, blood concentration) of a skin anomaly or normal tissue in a three dimensional map.
• This three dimensional map can be used to guide diagnostic or surgical
• Dangerous melanomas develop primarily by spreading in depth.
• the surface spread of the lesion is more easily measured, but not as useful for staging and prognosis as is the Breslow thickness which describes how deeply tumor cells have penetrated into the dermis.
• the Breslow thickess is prognostic factor in melanoma of the skin,
• NIR wavelengths penetrate more deeply because they are less likely to be absorbed/scattered. This provides one way to differentiate between superficial melanin absorption and deeper melanin absorption.
• By comparing the loss of polarization in light remitted from the tissue we have an indication of how deeply that light has penetrated the tissue. The more the polarization has changed from the original polarization, the more deeply the light has penetrated. Both these techniques [?] are used to determine the depth distribution of melanin. Variability in the depth distribution of the melanin across the lesion is a key indicator of melanoma.
• the spectral signature of the superficial layer as well as the deep layers of the skin can be determined.
• Spectrally characterizing the superficial layer will yield an estimation of melanin distribution in the superficial layer as shown in Fig. 5 (epidermal layer: 50-120 ⁇ depth).
• ⁇ ( 4 ( ⁇ )) the depth of the deep melanin at each x-y spatial coordinate can be estimated.
• volumetric irregularities can be compared as a strong indicator of melanocyte progression or abnormality which is a new capability.
• Fig. 7 shows top view of skin with vitiligo condition as an example of tissue with anomaly. Vitiligo has lower amount of melanin compare to normal skin therefore has lower optical attenuation.
• Fig. 8 shows the degree of linear polarization for both regions of vitiligo and normal skin condition.
• Fig. 9 shows top view of skin with mole condition as an example of tissue with anomaly. Mole has higher quantity of melanin compare to normal skin therefore has higher optical attenuation.
• Fig. 10 shows the degree of linear polarization for both regions of mole and normal skin condition.
• the disclosed technology can provide three dimensional reconstruction of tissue composition using a combination of hyperspectral, fluorescence and polarization based measurements.
• an attenuation correction is calculated using a ratio metric analysis of fluorescence anisotropy ("FA") and DLP to correct for attenuation-based artifacts.
• FA fluorescence anisotropy
• This calculation takes into account the effect of wavelength difference between excitation and emission wavelengths in DLP measurements by using a and ⁇ coefficients derived empirically for a particular tissue type or architecture, and a particular set of excitation and emission wavelengths.
• method 1 100 includes using at least two polarization modes of hyperspectral image cubes 1140 to calculate a three dimensional differential polarization data cube 1 160 using the following formula:
• ⁇ ( ⁇ ) ⁇ ⁇ ( ⁇ ) - ⁇ ( ⁇ )
• the three dimensional differential polarization data cube 660 is partially dependent on the surface reflection component of parallel polarization described in Fig. 1 and Fig. 2.
• Reflectance from the surface of the tissue is Lambertian in nature; that the amount of reflectance is proportional to the cosine of the angle of incidence of light encountering the tissue.
• Flat areas of the tissue appear bright while the areas with ridges and valleys become dark due to angle of incidence between the illuminating light and the tissue. Valleys in the tissue can act as light traps. Therefore, the image derived from the APOL data can transform to a map of tissue surface topography 1170.
• the measurements described herein provide a measurement space of light intensity as a function of position, wavelength and degree of polarization which can be transformed in a three dimensional array of voxels that characterize the tissue.
• the application of the method and system disclosed herein is not limited to the example described above.
• a system for capturing and processing multimode optical measurements is in Fig. 12.
• the system comprises an illumination beam path 1280 which presents illumination light to a sample 1240, an remitted light capture path 1290 that captures and detects light remitted by the sample 1240, and a control and data processing unit 1295 for controlling the illumination and detected light and processing the detected light.
• the illumination beam path 1280 comprises a light source 1210, an illumination spectral selection unit 1220, and an illumination polarization selection unit 1230.
• the remitted light capture path comprises an remitted light polarization selection unit 1250, and remitted light spectral selection unit 1260, and a detector 1270.
• the illumination light source 1210 may be at least one of a broadband lamp, such as tungsten or an arc lamp, a single wavelength laser, a multi-wavelength laser, a super continuum laser, a light emitting diode, or similar sources now or hereafter known in the art.
• the spectral selection units 1220 and 1260 may be an optical filter, an optical filter wheel, a diffraction grating, a liquid crystal tunable filter, an acousto-optic tunable filter, a plasmonic-based spectral selection device such as a metallic nanostructure, or similar spectral selection devices now or hereafter known in the art.
• the polarization selection units 1230 may be conventional polarizers such as rotatable crystal or wire grid polarizers or liquid crystal variable retarders, plasmonic metallic nanostructure based filters, or similar devices now or hereafter known in the art.
• the optical system 1280 may comprise free space optics, such as lenses, mirrors and prisms, fiber optics, integrated optics, liquid light guides, or other technology now or hereafter known in the art that can perform the same function.
• the illumination light source 1210 comprises a Xenon arc lamp incorporated in a spectral programmable light source, such as the product sold under the mark OneLight® Spectra by OneLight Corporation, Vancouver, BC, polarized in only one linear state.
• a spectral programmable light source such as the product sold under the mark OneLight® Spectra by OneLight Corporation, Vancouver, BC, polarized in only one linear state.
• the detected light from the tissue sample can be divided into two optical paths comprising cross and parallel polarizations using a beam-splitter and two orthogonally oriented polarizers and each polarization image detected by an individual CCD camera in each path, as will be understood by a person having ordinary skill in the art.
• the light remitted from the tissue sample may be spectrally filtered and passed through a polarization selection unit comprising a liquid crystal variable retarder and a linear polarizer that is oriented orthogonally to the illumination
• the liquid crystal variable retarder can be controlled to selectively rotate the polarization of the light remitted from the tissue sample prior to passing it through the linear polarizer, such that the fixed linear polarizer can act as a cross, 45 degree, parallel, or any other angle of polarization filter and the signal from each state can be sequentially captured with a single CCD camera.
• the system for acquiring the information may be deployed in an endoscopic measurement by delivering hyperspectral, polarized light though a light pipe or optical fiber, and receiving remitted light through the same or a
• Applicable polarization selection and spectral filtering methods may be selected by a person having ordinary skill in the art.
• the general goal of this disclosure is to arrive at an accurate three-dimensional representation of the structural characteristics of the tissue being tested based on multimode optical measurements.
• the disclosed system starts with the multimode measurements, produces values for the degree of linear polarization and fluorescence anisotropy as functions of wavelength ⁇ and position x',y' in measurement space, and estimates the structural characteristics of the tissue. Thereafter, the system corrects those characteristics based on comparisons of the predicted effect of the estimated structural characteristics on measured values or the linear polarization and fluorescence anisotropy. This is done by solving a multi-dimensional inverse problem as generally shown in Fig. 11.
• An estimation module 1180 produces tissue structural characteristics based on mul ti-mode optical measurements of tissue; that is, the principal inputs to the estimation T/US2014/014330
• the module are the degree of linear polarization as a function of wavelength and location of a detector element in measurement space, DLPx',y'( ), and the corrected fluorescent anisotropy as a function of wavelength and location of a detector element, FAx',y'( ), in measurement space, and produces as its output the structural characteristics of the tissue, such as the amount and location melanin in the tissue.
• the parameters of that model are estimated based on knowledge of the likely response of normal tissue to the illumination light that is to be used in the test and, if available, some understanding of the changes that might be caused by pathologies that may be present in the tissue. Those estimates are implemented by setting initial conditions for parameters of the estimation module.
• a three-dimensional forward model 1 190 is provided that predicts the optical response of tissue that should occur based on tissue structural characteristics, e.g., the amount and location of melanin in the tissue, and knowledge of the incident illumination light to be applied in the test, Jx,y(X)l 140; that is, the principal inputs of the forward model are structural characteristics of the tissue, and the principal outputs are the expected DLPx',y'( )l 150 and FAx',y'( )l 130.
• Another input to the forward model 1 190 is data representing the surface topography of the tissue produced by module 1170 in response to differences in the degree of linear polarization as a function of wavelength and the location of a detector element in measurement space, ADLPx',y'( ), computed by module 1 160.
• the tissue structural characteristics produced by the estimation module 1180 are provided as inputs to the forward model to produce as an output from the forward model the expected DLPx',y'(X)l 150 and FAx',y'(X)l 130 based on the known illumination light, the parameters of the inverse model and ADLPx ⁇ y'( )1160.
• the DLPx',y'( )1150 and FAx',y'(X)l 130 outputs produced by the forward model are then compared to the actual DLPx',y'(X) and FAx',y'( ) produced by
• DLPx ⁇ y'( ⁇ instruct)1150 and FAx ⁇ y'( )l 130 from the forward model 1190 are within
• the output of the system comprises the final structural tissue characteristics produced by the estimation module 1180.
• an inverse model algorithm implement by the foregoing modules and forward model wil l be applied to determine the tissue composition at each voxel.
• the forward model ⁇ 90 of the system may use various models for the propagation of light through tissue, such as the Monte Carlo method, diffusion theory, the random walk method, a radiative transfer model, or other similar models known in the art.
• tissue composition and optical properties including the known absorption, scattering and fluorescence properties of tissue.
• the output of any of the forward models is a data set corresponding to the data measured by measurement system 1200.
• the inverse problem algorithm postulates an initial state based on the standard or ideal forward model tissue composition values. It also postulates limits to the relative contribution of the tissue composition inputs that correspond to the real biological limits of the tissue composition.
• the inverse problem algorithm then iteratively adjusts the relative amounts of the tissue composition characteristics of the forward model until the output dataset and the measurement dataset converge.
• the limits to the relative tissue composition inputs constrain the iterations to stay within the bounds of biological plausibility and limit unnecessary calculations allowing the algorithm to converge faster and more efficiently.
• Fig. 13 illustrates a simplified system 1300 of three dimensional tissue characterizations using only hyperspectral and polarization based measurements. This includes using at least two polarization modes of hyperspectral data 1310 to create a degree of linear polarization spectral signature 1320. Exemplary polarization modes include linear polarized illumination and linear polarized detection in parallel, 45 degree, crossed or other orientations.
• the system 1300 may include the DLP spectral signature 1320 can be used to extract the initial estimate for three dimensional optical property of targeted tissue 1330.
• the three dimensional optical property of targeted tissue comprises at least two layers.
• the method 1300 may further include extracting the estimate of three dimensional composition and anatomical tissue mapping which can be used in three dimensional tissue models 1340.
• System 1300 solves an inverse problem in the same way as described with respect to system 1 100, and may use any of the forward models described above.
• Fig. 14 which illustrates a simplified system 1400 of three dimensional reconstruction of tissue composition using hyperspectral, fluorescence and polarization based measurements.
• This simplified system includes using at least two polarization modes of hyperspectral data 1440 to create a degree of linear polarization spectral signature 1450.
• Exemplary polarization modes include linear polarized illumination and linear polarized detection in parallel, 45 degree, parallel and other orientations.
• System 1400 may further include hyperspectral data in the form of hyperspectral image data.
• the hyperspectral image data may be structured in the form of a hyperspectral data cube comprising at least two polarization modes of fluorescence images 1410.
• the two polarization modes of fluorescent images 1410 can be analyzed to create fluorescence anisotropy mapping 1420.
• the method 1400 may further include attenuation correction of fluorescence anisotropy 1430 map using DLP mapping 1450 at the same wavelength range of corresponding fluorescence emission wavelength.
• System 1400 may include the DLP spectral signature 1450, and corrected fluorescence mapping 1430 can be used to extract the initial estimate for three
• the method 1400 may further include extracting the estimate of three dimensional composition and anatomical tissue mapping which can be used in three dimensional tissue models 1470.
• System 1400 solves an inverse problem in the same way as described with respect to system 1100, and may use any of the forward models described above.
• Embodiments of a method and subsystem for essentially eliminating the masking effect of superficial melanin and scattering are also disclosed herein.
• the method and system provide a polarized attenuation function APOL for more accurate skin 14 014330
• the method uses in vivo, non-invasive, hyperspectral, polarization sensitive imaging of skin based on specularly reflected and back-scattered light to determine anatomical and functional characteristics of skin with melanin or hemoglobin variations.
• the apparatus produces, and the method employs, two orthogonal, linear polarized hyperspectral image intensity datacubes.
• the method yields biologically plausible chromophore maps when applied to highly pigmented regions of skin.
• a dermoscope that enables in vivo, non-invasive polarized hyperspectral imaging of skin comprising a hyperspectral light source, polarization and other optics for illuminating a target and collecting remitted light, image detectors and control magnet analysis software that enables the multimode imaging-based measurement of skin lesions.
• the dermoscope that enables in vivo, non-invasive, polarized hyperspectral imaging of skin broadly comprises a console 1502 and a handpiece probe 1504.
• a computer in the console provides and controls the specimen illumination and data acquisition, image processing, archiving and data transmission.
• the illumination light is produced by a spectrally programmable OneLight® Spectra illumination system 1506 having a Xenon arc light source and microelectromechanically-based wavelength selection ability over the range from 468 nm to 857 nm.
• the console further comprises a computer 1507 and a display 1508 as well as appropriate input and output and data storage devices.
• the handpiece 1504 comprises two cameras; a beam splitter 1520; and fiber guides 1522 and 1524 that direct the light from the console illumination source to a fixture that positions this assembly at the correct depth to illuminate the tissue surface.
• the device preferably provides diffuse illumination to skin in a geometry that limits the amount of specular reflection to the detector.
• a ring-shaped linear polarizer 1526 is placed in front of the fiber optics to allow only linearly polarized light to illuminate the tissue surface.
• the two cameras share multi -element imaging optics 1528 and each camera has a polarization filter 1530 and 1532 respectively, which are oriented orthogonally to one another. This configuration captures images of the skin that maintain the linear polarization present in reflectance from both surface and deeper layers of tissue, and cross polarization images. Synchronized image acquisition by the two spatially registered cameras generates two images of an 11 mm ⁇ 16 mm area of skin in both parallel and cross polarizations. As shown in Fig.
• parallel polarized image stacks 1602 and cross polarized image stacks 1604 are acquired by hyperspectral imaging of the target area enabled by the sequential illumination with 33 wavelength bands from visible (468 nm) to near infrared range (857 ran), with a wavelength step interval of ⁇ 13 nm.
• Digital color images can be generated by programming the light source for broadband illumination to mimic typical Bayer filters that are used in conventional color cameras. These color images are provided for display or for comparison with standard dermoscopes. Additional system details for a specific such device are described in MacKinnon, N. B., et al. In vivo skin chromophore mapping using a multimode imaging dermoscope (SkinSpecTM), Proc. SPIE, 8587, 85870U (2013). In Fig. 16, the minimum spatially resolvable line-width detected by the P and X cameras was approximately 110 ⁇ , measured by imaging a USAF 1951 resolution test target.
• a calibration step is required to adjust the spatial and spectral intensity Z (x, y, ⁇ ), responses of the instrument, to correct for detector response, light source characteristics, and the instrument transfer functions.
• the imaging software determines camera exposure times for individual wavebands to optimize the cameras' dynamic range independent of illumination intensity variations.
• the calibration datacubes from imaging a SpectralonTM reflectance surface in both parallel (.3 ⁇ 4 and perpendicular ( x ) polarization states are computed using the following Eq.s: where 3 ⁇ 4 and are the reflectance measurements of skin hyperspectral images by parallel and cross polarized cameras, ⁇ g eeeraiea ar *d
• Z Sup4rrf ial is the reflectance of the skin superficial layer obtained by subtraction of the cross polarization image cube from the parallel polarization image cube.
• Eq. (2) shows how the calibration factor ( f xv, and scattering function (6, ⁇ ) at each pixel (x,y)and wavelength
• Z % includes superficially and deeply penetrating reflected light affected by both superficial and deep melanin as well as oxy- and deoxy- hemoglobin.
• OD log(Z (x, y, X)
• the minus sign is not required.
• Both Z g and ⁇ superficial include surface glare.
• C m _ s , C m _ d , C Hlf , and C oUb are the relative concentration of superficial and deep melanin, deoxy- and oxy hemoglobin, respectively;
• s m , s m , e olib are the absorption coefficients for melanin, deoxy-hemoglobin, and oxy-hemoglobin, respectively;
• L m _ x , L m ⁇ L Hh , and L Mb are the optical pathlength of superficial and deep melanin, deoxy- and oxy hemoglobin, respectively.
• AIN isolates the absorption of deep melanin, oxy- and deoxy hemoglobin thereby simplifying the quantification of these components.
• the term "deep” refers to light penetration into the reticular dermis to a depth of approximately 300 ⁇ or more.
• the pathlengths for the deep layer (dermis) are assumed to be equal for both deep melanin and hemoglobin ( L m _ d ⁇ L Hb «
• the cross polarization images show how the superficial and specular reflectance from the air-tissue interface is reduced and how more subsurface details (such as lesion boundary, micro-vascular patterns) become visible compared to the parallel polarization images.
• the cross-polarized optical density spectrum (OD L ) defined herein T/US2014/014330
• Figs. 17(c) and 17(g) next to the polarized attenuation spectrum shown in Figs. 17(d) and 17(h), respectively for melanocytic nevus and vitiligo.
• optical density spectra (OD ⁇ ) and the polarized attenuation spectra (A P0L ), as described in Eq. (5) are taken from three regions of interest: (central region) the melanocytic nevus core, (boundary region) halo, and surrounding normal skin.
• the optical density spectrum (OD x of the melanocytic nevus core (red square) shows the highest overall spectrum optical density (red line) due to its high melanin concentration.
• the relatively strong melanin contribution in the melanocytic nevus core results in a high polarized attenuation (A P0L ).
• An occlusion condition was induced by a plastic cuff on an imaged finger.
• a time sequence of 300* 150 pixels images from the same field of view at the dorsal side of the finger were cropped and concatenated to form a photographic strip chart shown, before putting on the cuff, during occlusion, and after removal of the cuff.
• the images were taken at thirty second intervals. The same experiment was repeated with the same subject's hand while probing the volar-side of the finger.
• Figs. 18(a) and 18(b) are color images of a portion of skin on the dorsal side of a subject's finger during application of occlusion captured under parallel and close polarization illumination respectively.
• Figs. 18(e) and 18(f) show color images of the volar side captured by parallel and cross polarization cameras, respectively.
• the volar side of the finger usually has less melanin concentration compared to the dorsal side of the finger.
• the darker color images during occlusion are caused by higher hemoglobin absorption due to more blood pooling in superficial blood vessels.
• Image contrast in cross polarization mode is enhanced due to the rejection of specular and superficial reflectance and preferentially selecting deeper penetrating light.
• Figs. 18(c) and 18(d) also show the optical density (OD A ) and polarized attenuation (A POL ) spectra from three representative images before, during, and after occlusion for the dorsal side of the finger.
• Figs. 18(g) and 18(h) show the optical density OD and polarized attention A POL for three corresponding images before, during and after occlusion for the volar side of the finger.
• Both OD ⁇ and A P0L spectra of both sides show higher attenuation during the occlusion period due to increased blood volume.
• the shape of attenuation spectra in the 500nm - 600 nra range more closely matches the single absorption peak of deoxy-hemoglobin absorption spectrum as compared to the two absorption peaks of oxy hemoglobin.
• This change in absorption trend is a result of progressive deoxygenation of the trapped blood due to the occlusion.
• the optical density spectra (OJD J and polarized attenuation spectra (A pQL ) show an increase in magnitude in the 500 run - 600 nm range.
• each solid line represents the mean of the corresponding pixel area (10x 10 pixels) shown in the related color images.
• the error bars represents the standard deviation of the attenuation at each wavelength for the pixels in the designated areas. While, the boxes in the color images appear to be from slightly different locations but are actually from the same anatomical location. The position change is due to slight movement of the finger during data acquisition.
• Oxy-hemoglobin has two absorption coefficient maxima at 542 run and 574 nm wavelengths and deoxy-hemoglobin exhibits a single absorption coefficient maximum at 545 nm.
• Melanin has a steadily linearly decreasing absorption trend in the spectral range from 600-700 nm and the slope of this curve increases proportional to the melanin content of an individual's skin Kollias, N, & Baqer, A., On the assessment of melanin in human skin in vivo, Photochem Photobiol, 43, 49-54 (1986). Light absorption by melanin and hemoglobin are similar in magnitude at wavelengths between 500 - 580 nm and hemoglobin or melanin concentration changes can be confused with one another during linear regression analysis.
• both oxy- and deoxy-hemoglobin absorption drops by one to two orders of magnitude at wavelengths longer than 600 nm, while the melanin absorption is still strong.
• the slope of the A POL function from 615 nm to 670 nm can be correlated with the concentration of deep melanin and is less affected by the influence of hemoglobin absorption. Therefore the deep melanin spatial distribution, Mel A (x,y ⁇ can be estimated as:
• the A PQi function can be corrected for the deep melanin absorption determined between 615 nm and 670 nm.
• the corrected spectrum A POL _ M)ll , orret . ttd can be analyzed to determine the oxy-and deoxy-hemoglobin concentrations using the linear least-square regression analysis in the 500 nm - 577 nm wavelength range (7 wavebands). This range encompasses the local absorption spectrum maxima of both oxy- and deoxy-hemoglobin.
• the resulting two-dimensional hemoglobin maps enable visualization of the superficial capillary network, as well as venous and arterial plexi, which are independent of melanin variations.
• Figs. 20(a) - 20(h) shows the derived chromophore maps of the skin with a melanocytic nevus Figs. 18(a) - 18(d) as well as skin with vitiligo Figs. 18(e) - 18(h).
• the skin melanin maps Figs. 18(b) and 18(f) were calculated from the optical density spectra ( D ) in cross-polarization mode.
• D optical density spectra
• a three- chromophore model was used, including melanin, oxy-hemoglobin and deoxy- hemoglobin employing curve-fitting algorithms with the extinction coefficients of the chromophores as primary vectors.
• Fig. 20(a) - 20(h) shows the derived chromophore maps of the skin with a melanocytic nevus Figs. 18(a) - 18(d) as well as skin with vitiligo Fig
• FIG. 18(c) shows how high melanin concentration is conducive to misestimation of the hemoglobin concentrations.
• the deep melanin estimation method described above was applied to correct this hemoglobin misestimation.
• Fig. 18(d) shows how this approach corrects the hemoglobin over- estimation in the nevus.
• the melanin corrected polarized attenuation spectrum (A -Mei corractad) was employed for hemoglobin estimation using a two-chromophore (oHb and Hb) model and curve-fitting algorithms with the extinction coefficients of oHb and Hb as primary vectors in the 500 run - 577 nm spectral range.
• Total hemoglobin was calculated by the summation of oxy-hemoglobin and deoxy-hemoglobin.
• the oxygenation saturation parameter (OSP) was calculated as a ratio of oxy-hemoglobin by the total hemoglobin as a percentage.
• Chromophore maps of skin with vitiligo were derived to evaluate the efficiency of the algorithm in skin tissue lacking melanin.
• the relative melanin distribution map for areas with vitiligo, Fig. 20(f) matched expectations for melanin.
• the estimated oxy- and deoxy-hemoglobin shows high correlation with melanin in vitiligo, the same effect shown in highly pigmented nevus, Fig 20(g).
• the same melanin correction method to the A POL spectrum, the same correction effect in the areas with vitiligo was confirmed, providing a more biologically plausible hemoglobin distribution, Fig. 18(h).
• the melanocyte nevus is more strongly affected due to the greater melanin difference to the surrounding normal skin then when the correction is observed for the vitiligo condition.
• Fig. 21 compares the cross-polarized color images and skin chromophore map sets for the dorsal-side and volar-side of a human finger during venous occlusion.
• the melanin corrected polarized attenuation spectra was from a region of interest (100 x 150 pixels) and fitted to a two chromophore skin model (oHb and Hb) in the 500 nm - 577 nm range.
• the deep melanin estimation was the same method that was presented for nevus and vitiligo described in a previous section of this manuscript.
• Visible reflectance hyperspectral imaging characterization of a noninvasive, in vivo system for determining tissue perfusion, Ana! Chem, 74, 2021-2028 (2002) Matthijs, D., Hondebrink, E., van Leeuwen, T., & Steenbergen, W. Time domain algorithm for accelerated determination of the first order moment of photo current fluctuations in high speed laser Doppler perfusion imaging, Med Bio Eng Co p, 47, 1 103-1 109 (2009).
• occlusion measurements three measurements of the subject's finger were initiated before initiating occlusion (by a plastic cuff on subject's finger). Five post- occlusion measurements were taken, then another five measurements after cuff removal (during reperfusion). AH data were taken at 30 second intervals. Two sets of measurements, one from the volar surface of the finger and the other from the dorsal surface of the finger were acquired. This permitted a comparison of the effect of melanin change on tissue oxygenation estimation as the volar side of the finger had less melanin. Vyas, S., Banerjee, A., & Burlina, P. Estimating physiological skin parameters from hyperspectral signatures. , JBiomed Opt, 18, 057008 (2013).

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